High resolution dynamic detector for imaging and dosimetry in megavoltage radiation therapy

ABSTRACT

Disclosed herein are variations of megavoltage (MV) detectors that may be used for acquiring high resolution dynamic images and dose measurements in patients. One variation of a MV detector comprises a scintillating optical fiber plate, a photodiode array configured to receive light data from the optical fibers, and readout electronics. In some variations, the scintillating optical fiber plate comprises one or more fibers that are focused to the radiation source. The diameters of the fibers may be smaller than the pixels of the photodiode array. In some variations, the fiber diameter is on the order of about 2 to about 100 times smaller than the width of a photodiode array pixel, e.g., about 20 times smaller. Also disclosed herein are methods of manufacturing a focused scintillating fiber optic plate.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application is a continuation of U.S. application Ser. No.17/711,606, filed Apr. 1, 2022, which is a continuation of U.S.application Ser. No. 16/993,029, filed Aug. 13, 2020, now U.S. Pat. No.11,300,692, which is a divisional of U.S. application Ser. No.15/921,219, filed on Mar. 14, 2018, now U.S. Pat. No. 10,775,517, whichis a continuation of International Application No. PCT/US2016/051750,filed on Sep. 14, 2016, which claims priority to U.S. Provisional PatentApplication No. 62/218,425, filed on Sep. 14, 2015, each of which arehereby incorporated by reference in their entirety.

BACKGROUND

In modern megavoltage (MV) X-ray radiation therapy, the radiationdetector used to determine the position of the patient within thetherapy system is typically different from the detector used to measurethe radiation dose delivered to that patient during the therapy session.One reason that two separate detectors are used is that the sensitivityof the detector for anatomical imaging is different from the sensitivityof the detector for measuring radiation dosage. Detectors suitable foranatomical imaging, such as electronic portal imaging detectors (EPID),often use scintillating materials made of heavy elements, such asGadolinium oxysulfide (GOS) or cesium iodide doped with thallium(CsI(Tl)). However, because such materials have a radiation doseresponse which is vastly different from tissue or water, they are notsuitable for patient dosimetry measurements. Detectors suitable formeasuring the radiation dose delivered to a patient are typically madeof a low-Z material, which have a similar radiation dose response tothat of tissue or water. Examples may include dry air in ion chambers,diamond, silicon, etc. Such detectors are not appropriate for acquiringimages because of their low stopping power for X-rays. Althoughincreasing the thickness of the low-Z material may allow for the captureof more X-rays, the resultant image may have poor spatial resolution.

Accordingly, a radiation detection system that is capable of bothacquiring images at a useful spatial resolution and measuring theradiation dose delivered to a patient is desirable.

BRIEF SUMMARY

Disclosed herein is a megavoltage (MV) detector that may be used foracquiring high resolution dynamic images and dose measurements inpatients. One variation of a MV detector may comprise a light photonconverter, a light sensor array (e.g., a photodiode array), and opticalfibers optically connecting the photon converter and photodiode array.In some variations, a MV detector may comprise optical fibers configuredto convert X-rays into photons (e.g., scintillating optical fibers) andact as a light guide or channel to deliver the photons to a light sensorarray. For example, a MV detector may comprise a scintillating fiberoptic plate coupled to an array of photodiodes and thin-film transistors(TFT), where each photodiode and TFT pair represents a pixel of animage. The scintillating fiber optic plate may comprise a plurality ofmicron-sized optical fibers bundled together, and the fibers may beangled such that they are aligned along the rays of an X-ray sourcelocated across from the MV detector (i.e., the fibers of thescintillating fiber optic plate are “focused” to the X-ray source). Thefibers may be made of a low-Z material (e.g., plastic). In somevariations, the fibers of a focused scintillating fiber optic plate(FSFOP) may be clustered into groups, for example, a left cluster (orblock, or module) of fibers and a right cluster (or block or module) offibers. FSFOPs may have a thickness, which may correspond to the lengthof the fibers, from about 1.5 cm to about 30 cm, e.g., about 5 cm thick,e.g., about 4 cm thick. The diameters of the fibers may be smaller thanthe pixels of the photodiode array. In some variations, the fiberdiameter may be on the order of about 2 to about 100 times smaller thanthe width of a photodiode array pixel, e.g., about 20 times smaller.Also disclosed herein are methods of manufacturing MV detectors thatcomprise a FSFOP and a photodiode/TFT array and methods of using the MVdetectors described herein to both generate a patient image and todetermine the amount and location of radiation (i.e., the radiationdose) emitted by a source and/or applied to the patient during aradiation therapy treatment session.

One variation of a radiation detector may comprise a fiber optic arraycomprising a plurality of scintillating fibers, each fiber having adiameter, an input face, an output face, and a longitudinal axistherebetween, and a photodiode array coupled to the output faces of thefibers in the array. The fibers may be focused to a radiation source.The photodiode array may comprise a plurality of photodiodesrepresenting a plurality of pixels, where each pixel has a pixel width.The fiber diameter may be smaller than the pixel width and a pluralityof the output faces of a plurality fibers may be in contact with eachpixel of the photodiode array. In some variations, each pixel of thephotodiode array may directly contact a plurality of fibers. Thelongitudinal axis of each fiber may be aligned with a propagation axisof a ray of radiation emanating from the radiation source, and/or thelongitudinal axes of the fibers may be aligned to rays of a radiationbeam emitted by the radiation source. The fiber optic array may have athickness from about 1.5 cm to about 5 cm. The ratio of the fiberdiameter to the pixel width may be from about 1:10 to about 1:100. Thefiber diameter may be from about 5 μm to about 10 μm (e.g., about 10μm), and the pixel width may be from about 150 μm to about 1000 μm(e.g., about 400 μm). The scintillating fibers may comprise a materialhaving a density similar to that of water, e.g., about 1 g/cm³, or aboutabout 1.18 g/cm³. For example, the material may be a plastic. In somevariations, the plurality of scintillating fibers may be clustered inblocks, where the fibers of each fiber block may be aligned to a uniqueor different portion of the radiation source beam. The radiation sourcemay emit a fan beam having a right portion and a left portion, and thefiber optic array may comprise a first block of scintillating fibersaligned toward the right portion of the fan beam and a second block ofscintillating fibers aligned toward the left portion of the fan beam. Insome variations, a radiation detector may comprise a sheet of metalbetween each of the blocks. The metal may comprise a high-Z metal, andmay have a thickness from about 0.1 mm to about 2 mm.

In some variations, the fiber optic array may comprise a top surfacehaving a first surface area and a bottom surface having a second surfacearea, where the first surface area is less than the second surface areaand the bottom surface contacts the photodiode array. Optionally, ametal sheet may be disposed over the top surface of a fiber optic array.The metal sheet may be a low-Z metal, such as copper or aluminum. Themetal sheet may have a thickness from about 0.05 mm to about 1 mm, e.g.,about 0.2 mm. Alternatively or additionally, a fiber optic array maycomprise a top surface and a bottom surface that contacts the photodiodearray, where the top surface may be coated with one of alight-reflective paint or a light-absorbing paint.

Disclosed herein is one variation of a method of manufacturing anoptical fiber plate for a detector. One method may comprise providing ablock of parallel scintillating optical fibers, slicing the block alonga first axis that transects the optical fibers at a first angle withrespect to the parallel fibers to create a top surface, slicing theblock along a second axis that transects the optical fibers at a firstdistance away from the top surface to create a bottom surface, whereinthe bottom surface is parallel to the top surface, slicing the blockalong a third axis to create a first side surface that is at a secondangle with respect to the top surface, slicing the block along a fourthaxis at a second distance away from the third axis to create a secondside surface, wherein the second side surface is at a third angle withrespect to the top surface. The first distance may correspond to athickness of an optical fiber plate and the second distance maycorrespond to a width of the optical fiber plate. The third angle may bedifferent from the second angle, and/or the first angle may be fromabout 10 degrees to about 180 degrees, and/or the second and thirdangles are from about 0.5 degrees to about 40 degrees. The scintillatingfiber located at the center of the optical fiber plate may be at afourth angle such that the center scintillating fiber is aligned with aray of a radiation source fan beam located at a fixed distance away fromthe optical fiber plate. The method may further comprise comprisingslicing the block along a fifth axis to create a third side surface anda sixth axis at a third distance away from the fifth axis to create afourth side surface, where the third side surface is at a fourth anglewith respect to the top surface and the fourth side surface is at afifth angle with respect to the top surface. The scintillating opticalfibers may comprise plastic scintillating optical fibers, and/or mayhave diameters from about 5 μm to about 10 μm.

Another method of manufacturing an optical fiber plate may compriseproviding a billet of tapered scintillating optical fibers, slicing thebillet along a first axis that transects all of the fibers to create atop surface, slicing the billet along a second axis that transects allof the fibers at a first distance away from the first cut to create abottom surface, wherein the bottom surface is parallel to the topsurface, slicing the billet along a third axis to create a first sidesurface that is at a first angle with respect to the top surface, andslicing the billet along a fourth axis at a second distance away fromthe third axis to create a second side surface that is at a second anglewith respect to the top surface. The billet of scintillating opticalfibers may be thermally tapered. The taper angle of the billet ofoptical fibers may be determined at least in part by the shape of aradiation beam from a radiation source that is to be located at a fixeddistance away from the optical fiber plate. The first side surface mayextend along a length of a first boundary fiber in the tapered billet,and the second side surface may extend along a length of a secondboundary fiber in the tapered billet. The scintillating optical fibersmay comprise plastic scintillating optical fibers, and/or may havediameters from about 5 μm to about 10 μm.

Another variation of a radiation detector may comprise an optical fiberplate comprising a matrix of optical fiber modules, wherein each opticalfiber module comprises a plurality of scintillating optical fibersincluding a central scintillating optical fiber located in the center ofeach fiber module, and a photodiode array coupled to a bottom surface ofthe optical fiber plate, the photodiode array comprising a plurality ofphotodiodes representing a plurality of pixels, each pixel having apixel width. The central scintillating optical fiber of each of theoptical fiber modules may be focused to a radiation source to be locatedat a fixed distance away from the radiation detector. The radiationdetector may optionally comprise opaque septa between each of theoptical fiber modules. The longitudinal axis of the centralscintillating optical fiber of each of the optical fiber modules may bealigned along a ray of a radiation beam generated by the radiationsource. A longitudinal axis of each of the optical fibers in each of themodules may be aligned along different rays of the radiation beam.Alternatively or additionally, the longitudinal axis of the centralscintillating fiber of each of the plurality of fiber modules may bealigned with a different ray of the radiation beam. The optical fiberplate has a thickness from about 1.5 cm to about 5 cm. The optical fiberplate further comprises a low-Z metal plate disposed over a top surfaceof the plate. In some variations, the area of the bottom surface of theoptical fiber plate may be greater than the area of a top surface of theoptical fiber plate. The fiber diameter may be smaller than the pixelwidth and a plurality of output faces of the fibers may be in contactwith each pixel of the photodiode array. The scintillating opticalfibers may comprise plastic scintillating optical fibers. Thescintillating optical fibers have a diameter from about 5 μm to about 10μm. In some variations, the scintillating optical fibers may comprisematerials having different refractive indices.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1A is an end view schematic depiction of one variation of aradiation therapy system. FIG. 1B is a side view of a portion of thesystem of FIG. 1A. FIG. 1C is an end view schematic depiction of anothervariation of a radiation therapy system comprising another variation ofa MV detector. FIG. 1D is a schematic depiction of focused fibers. FIG.1E is a schematic depiction of unfocused fibers. FIG. 1F depictspercentage dose deposition (PDD) curves for a water-like material fordifferent radiation source (e.g., X-ray) energies, including a PDD curvefor cobalt for comparison. FIG. 1G is a plot of detection efficiency asa function of detector thickness over several radiation source (e.g.,X-ray) energies, where the detector comprises a water-like material.

FIG. 2 depicts a cross-sectional view of one variation of a fiber opticplate and flat panel photodiode/TFT array of a MV detector.

FIG. 3 depicts a perspective view of one variation of a fiber opticplate.

FIG. 4 is a representation of the mapping between the fibers of onevariation of a fiber optic plate and the pixel boundaries of aphotodiode/TFT array.

FIG. 5 is a schematic representation of one variation of photodiodearray and the readout electronics.

FIG. 6 is a timing diagram of one variation of a treatment imagesequencing method.

FIG. 7A is a perspective view of one fiber optic block of a fiber opticplate of a MV detector. FIG. 7B is a side schematic view of one fiberoptic block made in accordance with the method outlined in the flowchartof FIG. 7C.

FIG. 8A is a side schematic view of one fiber optic block made inaccordance with the method outlined in the flowchart of FIG. 8B. FIG. 8Cdepicts the coordinates for the computation of fiber tilt angles. FIG.8D is a perspective view of the fiber optic block of FIG. 8A.

FIGS. 9A and 9B depict a first side schematic view of one variation of afiber optic plate (along the X-direction) and FIG. 9C depicts Table 1,which summarizes the dimensions of the fiber optic plate depicted inFIGS. 9A and 9B.

FIG. 10A depicts a second side schematic view (along the Z-direction) ofthe fiber optic plate of FIG. 9A. FIG. 10B depicts Table 2, whichsummarizes the dimensions of a portion of one variation of a fiber opticplate depicted in FIG. 10A.

FIG. 11A depicts Table 3, which summarizes example parameters (e.g.,rates of rotation and movement) for one variation of a treatmentsession. FIG. 11B schematically depicts entrance beams and exit beamsfor a plurality of gantry or firing angles. FIG. 11C is a flowchartdepiction of one method for determining a dose distribution map or imageusing the MV detectors described herein.

FIG. 12A is a graphical representation the same angle projection imagingof a moving fiducial marker (triangle) across the field of view, or theslice width, of an X-ray beam generated by the linac. FIG. 12B is ashift-and-add graphical representation of one variation of an image anddose averaging scheme using the movement of the fiducial marker of FIG.12A as an example.

FIG. 13 is a graphical representation of the number of images availablefor averaging as a function of couch displacement for one variation of aradiation therapy system.

DETAILED DESCRIPTION

A radiation therapy system may comprise a patient area, a radiationsource located on one side of the patient area configured to applyradiation to a patient, and a MV detector located on a second side ofthe patient area (e.g., generally opposite the radiation source)configured to detect radiation from the source. The readings from the MVdetector may be used to generate an image of a patient, and/or may beused to compute the radiation dose delivered to the patient from theradiation source. However, MV detectors that are suitable for detectingradiation for the purpose of generating a high-resolution anatomicalimage are not usually suitable for measuring the radiation dose appliedto the patient. For example, an MV detector for anatomical imaging maycomprise a layer of hi-Z material, but since a hi-Z material has aradiation dose response that is different from the radiation doseresponse of tissue or water, it may be computationally challenging orintensive to determine the dose delivered to the patient. MV detectorssuitable for measuring radiation dosage are usually not suitable forgenerating high resolution anatomical images. MV detectors for dosimetrypurposes usually comprise a layer of low-Z material (e.g., having asimilar or equivalent response to human tissue and/or water) whichfacilitates the radiation dose computations, but does not provide enoughstopping power to capture X-rays for anatomical imaging.

Disclosed herein are novel MV detectors for high resolution dynamicimaging and dose measurement. These MV detectors may be used in aradiation therapy system in order to provide the practitioner withinformation about the anatomy and/or position of the patient, as well asthe radiation dose delivered to a particular location in the patientduring a radiotherapy treatment session. Also disclosed herein aremethods of manufacturing such MV detectors. One variation of a detectormay comprise a fiber optic plate, a light sensor array such as aphotodiode array (i.e., a photodiode/TFT array), and readoutelectronics. The fiber optic plate may comprise a plurality ofscintillating optical fibers bundled together and oriented such that thelongitudinal axes of the fibers are pointed towards a radiation sourcethat is to be located across the detector in the radiation therapysystem (i.e., the fibers are focused to the radiation source). Thefibers may be made of plastic, and/or any material(s) having X-raylinear attenuation coefficients similar to that of water or humantissue. The linear attenuation coefficient of such materials may beabout 1 g/cm³ for water and plastics. The fibers may be focused to aradiation source (i.e., their longitudinal axes may be aligned with therays of a radiation beam produced by the radiation source). Thethickness of the fiber optic plate may be from about 1.5 cm to about 20cm depending on the X-ray energy of the radiation source and therequirement for the detective quantum efficiency (DQE), e.g., about 5cm, about 3 cm to about 5 cm, about 5 cm to about 10 cm, about 7 cm toabout 12 cm, about 10 cm to about 20 cm, etc. The thickness of the fiberoptic plate may also be selected such that it may provide sufficientstopping power for detecting 6MV X-rays for imaging. Orienting thefibers in the optic plate such that their longitudinal axes lie alongthe ray lines of a radiation beam may help to preserve data relating tothe directionality of the radiation beam incident upon the fiber opticplate, and/or reduce light scatter between the fibers and which may helpto preserve the pixel resolution of an anatomical image.

FIGS. 1A-B depict one variation of a radiation therapy system. Aradiation therapy system 100 may comprise a gantry 102 configured torotate around a patient area 104, a radiation source 106 coupled to thegantry, a MV detector 108 coupled to the gantry such that it is locatedacross from (i.e., opposite to) the radiation source, and a controller(not shown) in communication with the gantry, radiation source and theMV detector. The gantry 102 may be a rotatable ring or circular gantryand may have a plurality of firing locations at differentcircumferential locations or angles on the gantry. A collimator 107 maybe disposed in front of the radiation source 106 (i.e., within the beampath). A movable patient couch 110 may be located within the patientarea 104, and may be configured to move the patient 112 within thepatient area, relative to the radiation beam 114 (e.g., fan beam)applied by the radiation source 106. The patient couch may be configuredto move the patient in a direction that is orthogonal to a plane ofradiation emitted by the radiation source 106, for example, into and outof the patient area 104 which may be defined by the bore or centrallumen of the gantry (as depicted by the side view of FIG. 1B). Thegantry may be a circular gantry, and may be an open or closed boregantry, etc. The location of the radiation source 106 may be fixedrelative to the location of the MV detector 108. For example, for aparticular radiation therapy system or a particular radiation therapysession, the distance between the radiation source and the MV detector,as well as their locations relative to each other, may be determinedduring and/or before manufacturing, and/or before the therapy session.Maintaining a fixed relative location may allow a MV detector to betailored to the beam spread of the radiation source (e.g., focused tothe radiation source), which may help to reduce the amount of noisedetected and/or preserve pixel resolution.

One variation of a MV detector that may be included in a radiotherapysystem may comprise a focused scintillating fiber optic plate (FSFOP), alight sensor array such as a photodiode array (e.g., an amorphoussilicon (a-Si) photodiode/TFT array) which may be optically coupled tothe fiber optic plate, and high frame rate readout electronics incommunication with the light sensor array. The fiber optic plate may bean optical device comprising a plurality (e.g., bundle, group, cluster,or block) of micron-sized optical fibers. The optical fibers mayscintillate in response to X-ray photons, and may be used both as anX-ray-to-light photon converter and as a light guide to transport lightphotons to the light sensor array. Generating visible light photons inan optical fiber and transporting those light photons within the samefiber may help to reduce the spread of the light photons along thelength of the fibers. Each pixel of the light sensor array may contactand/or receive light photons from a plurality of optical fibers. Thefibers in the fiber optic plate may be focused such that they point tothe radiation source that is located across from the MV detector in theradiation therapy system.

As described throughout this document, “focused” optical fibers refer toalignment of the longitudinal axes of the optical fibers to the rays ofa radiation beam from a radiation source that is to be located acrossfrom the MV detector in a radiation therapy or imaging system. Forexample, if a radiation source is represented as a focal spot, theradiation beam(s) emanating from it may be represented by a plurality ofrays that intersect with and/or originate from that focal spot. Thegeometry of the radiation beam emanating from the focal spot may beshaped as a cone, a fan, or any other shape as may be determined by oneor more radiation beam collimators and/or jaws. An optical fiber may bedescribed as “focused” if its longitudinal axis is approximately orsubstantially or exactly aligned with a ray of the radiation beam. Thatis, if the longitudinal axis or length of a focused optical fiber isextrapolated or extended, it would intersect with the focal spot of theradiation source, or nearly or substantially intersect with the focalspot. A cluster, block, or module of optical fibers may be described as“focused” if a majority of the fibers in the cluster, block, or moduleare aligned with one or more radiation rays (e.g., about 50% to about100%, about 50%, about 60%, about 70%, about 80%, about 90%, about 75%,etc.). FIG. 1D schematically depicts examples of individual opticalfibers 142 a, 142 b, 142 c, 142 d, 142 e that are focused to a radiationsource 140 (note that the fibers are not drawn to scale and may havewidths and lengths that differ from the depiction in FIG. 1D). Theradiation source 140 may emit a radiation beam 144 that may comprise aplurality of rays 146. The longitudinal axes 148 a-148 e of each of thefibers 142 a-142 e may be aligned to a ray 146 of the emitted beam 144.In contrast, FIG. 1E schematically depicts examples of individualoptical fibers that are not focused to a radiation source. In FIG. 1E,the longitudinal axes 158 a-158 e of most of the fibers 152 a-152 e arenot aligned to a ray of the emitted beam. In this variation, the onlyfiber that may be considered focused is the central fiber 152 c that isdirectly across from the radiation source. Alternatively oradditionally, in variations of a MV detector comprising a focused fiberoptic plate comprising a plurality of clusters, blocks, or modules, asingle cluster, block, or module comprising a plurality of opticalfibers may be described as “focused” if the center-most optical fiber ofthat cluster, block or module is substantially or approximately orexactly aligned with a ray of the radiation beam, and the other fibersin the cluster, block, or module are substantially parallel to thecenter-most fiber. In the description of a focused fiber optic plateand/or a focused optical fiber, it is understood that the plate and/orfiber is focused to a radiation source, even if a radiation source isnot explicitly described or depicted in the description.

While FIGS. 1A and 1B depict a single continuous large rectangularshaped MV detector that spans the full beam width (e.g., the spread orillumination field of the radiation beam at a particular distance fromthe X-ray or radiation source, the region of irradiation at a particulardistance from the radiation source), in other variations, a MV detectormay be divided into multiple rectangular or square modules or blocks.For example, as depicted in FIG. 1C, the modules or blocks may beassembled or tiled together to form an overall arc shape that is focusedto the X-ray source. In some variations, the MV detectors may bearranged along the curvature of the gantry, i.e., the curve of the arcof modules, may have the same radius of curvature as a circle that isconcentric with the gantry. In some variations, the radius of curvatureof the overall MV detector may be different from (e.g., greater than orless than) the radius of curvature of the gantry. In some variations,the curvature of the arc shape may correspond to the angular spread ofthe radiation beam (e.g., fan beam) of the X-ray source. Depending onthe requirement for focusing accuracy, individual fibers in the modulesmay be focused to the X-ray source or clustered into blocks where thefibers are parallel to each other, as described further below.

The overall shape and size of the MV detector may vary as may bedesirable according to different X-ray beam geometries, gantrygeometries (e.g., size, shape), and/or to fit with the particulararrangement of different radiation treatment systems. For example, aC-arm radiation therapy system may comprise a square-shaped MV detector,while a helical tomotherapy system may comprise a long, rectangulardetector. Although the examples of MV detectors described herein are inthe context of a helical tomotherapy radiation system, it should beunderstood that the shape and size of the MV detectors may varydepending on the arrangement of the radiation therapy system desired.The orientation of the fibers in the fiber optic plate may be adjustedin accordance with the relative locations of the MV detector and theradiation source of a particular radiation therapy system such that theoptical fibers of the MV detector are focused to the radiation source.

Scintillating fibers (which may be, in some variations, made of plastic)may be bundled together to form a thickened fiber optic plate. As thethickness of the fiber optic plate increases, the X-ray stopping powerof the plate may also increase. Scintillating fibers emit visiblephotons when irradiated by X-rays. The scintillating fibers of the fiberoptic plate may be arranged such that they are focused to the radiationsource (e.g., to a focal spot of a linac). As described above, the FSFOPmay function both as an X-ray-to-optical photon converter and as a lightchannel to guide the optical photons to the photodiode/TFT array. Theoptical fibers of the plate have a relatively small diameter as comparedto the size of the pixels of the light sensor array (e.g., aphotodiode/TFT array). For example, the diameter of a scintillatingoptical fiber may be from about 1 μm to about 50 μm (e.g., from about 5μm to about 50 μm), while the width of a pixel may be from about 100 μmto about 1000 μm, e.g., about 100 μm to about 200 μm, about 400 μm. Theratio of the fiber diameter to pixel width may be from about 1:5 toabout 1:100, e.g., about 1:10, about 1:15, about 1:20, about 1:30, etc.The small diameter and the light channeling property of thescintillating optical fibers may address the factors that hinder pixelresolution and image clarity. That is, fiber optic plates that havefibers with diameters smaller than the pixel size of a photodiode/TFTarray may give rise to images with better resolution than fiber opticplates that have larger diameter fibers (e.g., where the fiber diameteris approximately the same as the pixel size). Fiber optic plates wherethe fibers have a diameter that is on the same order of magnitude (orsubstantially similar to) the width/length of a pixel may require thateach fiber is aligned with exactly one pixel to facilitate the formationof images with a desirable resolution and sharpness. Fiber misalignmentmay cause the visible light photons from a single fiber to be randomlyfed to two or more pixels, which may contribute to image blur and lossof resolution. However, precise and specific alignment between thefibers of a FSFOP and the pixels of a light sensor array may bechallenging due to imprecise positioning of the FSFOP and the lightsensor array during the manufacturing process. Misalignment betweenfibers and pixels may cause cross-talk between neighboring pixels thatare receiving light input from the same fiber. In variations where thefibers have a diameter that is approximately the same size as the pixelwidth of the photodiode/TFT array, such cross-talk may degrade both thespatial resolution and the signal uniformity of the photodiode/TFTarray. For example, if the size of a pixel of a photodiode/TFT array isabout 400 μm and a fiber has a diameter of the same or similar size,misalignment between the fiber and the pixel by an offset of about 200μm may cause the light signal from the fiber to be input to twoneighboring pixels equally. In this hypothetical scenario, the spatialresolution of the photodiode/TFT array may degrade up to about 800 μm,despite the native pixel resolution being about 400 μm. The degradationto pixel resolution may be reduced by using fibers having a much smallerdiameter than the pixel size. For example, where the fiber diameter isabout 50 μm and the photodiode/TFT array pixel size is about 400 μm, asingle fiber that contacts two neighboring pixels (i.e., providing lightinput to two adjacent pixels) would result in a light signal spread ofno more than about its diameter of 50 μm (instead of about 400 μm in theprevious example). This reduced level of cross-talk may help reduce thedegradation of spatial resolution of the photodiode/TFT array, and insome variations, limit the level of cross-talk between pixels such thatthe degradation to spatial resolution is negligible.

Misalignment between fibers and pixels of approximately the same sizemay cause the septa between the fibers to occupy light-sensitive regionsof the photodiode/TFT array. The thickness of septa or cladding inbetween each of the fibers may result in relatively large and contiguousareas on the sensor pixel where light data may not be sensed (e.g.,sensor “dead zones”). Reducing the diameter of the fibers such that theyare substantially smaller than the size of a pixel may help to alleviatethese issues. Reducing the diameter of the fibers such that they aresubstantially smaller than the size of a pixel may help to alleviatethese issues. Fiber optic plates where the fiber diameters aresubstantially smaller than the size of a pixel (e.g., where the fiberdiameter is at least about 4 times, or about 5 times, or about 8 times,or about 10 times, or about 13 times, or about 15 times, or about 25times, or about 33 times, or about 50 times, or about 75 times, or about100 times, etc. smaller than the smallest dimension of a pixel) may notrequire a one-to-one fiber-to-pixel mapping in order to preserve imagequality. Rather, the light output of a plurality of fibers may be mappedto a single pixel. Since the longitudinal axes of a population of fibersmay be, as a group, aligned to the rays of a radiation beam emitted by aradiation source (i.e., focused to a radiation source), the averagelight signal acquired by the photodiode/TFT array from this populationmay preserve and/or approximate the direction from which the incidentX-ray originated. As a result, averaging the signals from multiplefibers across a single pixel may help to smooth out an image.Furthermore, the septa or cladding between each of the smaller-diameterfibers may be distributed across the pixel, instead of in a fewcontiguous regions, which may help to reduce the size of individual deadzones and average the effect of such dead zones across the area of thepixels. In some variations, averaging signals across multiple pixels mayalso help to smooth out the image.

The fibers of a fiber optic plate (e.g., a FSFOP) may be made of anymaterial(s) that have mass-energy absorption coefficients and collisionstopping powers similar to that of water. For example, the fibers may bemade of one or more low-Z, low-density materials, and/or materials withlow atomic numbers that are similar to water, and/or any materials witha water-equivalent dose response to radiation. Plastic materials such aspolystyrene (e.g., BCF-60), and acrylic may be included in the fibers ofthe fiber optic plate. Measurement of the amount of radiation deliveredto the scintillating optical fibers that have a dose response similar towater may then be used as a model or proxy for the amount of radiationthat may be delivered to tissue (e.g., a region of interest in apatient). MV detectors comprising scintillating optical fibers having adose response similar to water may also allow for calibrations (e.g.,dose calibrations) without a water phantom. A dose-depth distributioncurve, such as a percentage dose deposition (PDD) curve for a variety ofX-ray source energy levels (e.g., 4 MeV, 6 MeV, 12 MeV, 25 MeV) as afunction of water depth is depicted in FIG. 1F. For comparison purposes,the PDD curve for cobalt for a 1.25 MeV energy X-ray source is alsodepicted. For the purposes of this document, the power or energy levelof the X-ray radiation emitted by a radiation source will be describedas either “MV” (megavolts) or “MeV” (mega electronvolts), where thesequantities are related as follows: an electron moving through anelectric potential difference of 1 MV gains 1 MeV of energy. Thecumulative amount of radiation dose applied to a tissue across aparticular depth may be estimated by integrating the dose-depth curveacross the length of the fiber (which maps to the thickness of the FOP).The PDD for water is relatively small at superficial skin depths (e.g.,at depths less than about 0.5 cm) and increases to peak values at depthsof about 1.4 cm for a 4 MeV beam and about 3.5 cm for a 25 MeV beam. Assuch, the thickness of a FSFOP, or the length of the fibers of a FSFOP,may be from approximately 1.5 cm (or at least about 1.5 cm) toapproximately 4 cm (or approximately 5 cm) in order to absorb and/ormeasure the peak dose delivered by a X-ray source having an energy fromabout 6 MeV to about 25 MeV (which may provide an estimate of theradiation dose delivered to a patient). A FSFOP thickness (or fiberlength) that is shorter than 1.5 cm may not allow the detector tocapture the peak dose delivered, may result in a measurement that is asignificant underestimate of the amount of radiation that would bedelivered to a region of interest in a patient (which is usually morethan 1.5 cm beneath the skin surface). Accordingly, an MV detectoradapted for dosimetry purposes may comprise an FSFOP where the fibermaterial(s) have a water-equivalent dose response, as described above,and have a plate thickness (i.e., fiber length) of at least about 1.5 cmin order to capture at least the peak radiation dose. In principle,thickening the FSFOP (i.e., lengthening the fibers) may increase theefficiency of the MV detector, however, increasing the thickness (orlength) may introduce other challenges, such as difficulty in aligningthe longitudinal axis of a lengthy fiber to the rays of a radiation beamemitted by an X-ray or radiation source, which may result in lightspread (e.g., blurring), etc. While the thickness of a FSFOP (e.g.,lengths of the fibers) may be longer than about 1.5 cm (e.g., about 2cm, about 2.5 cm, about 3 cm, about 3.5 cm, about 4 cm, about 4.5 cm,about 5 cm, about 5.5 cm, about 6 cm, about 6.5 cm, about 7 cm, about7.5 cm, about 8 cm, about 8.5 cm, about 9 cm, about 9.5 cm, about 10 cm,etc.), the desired thickness of the FSFOP may be determined at least inpart by one or more of the following: the physical space limitations ofthe radiation therapy system, desired X-ray stopping power (andtherefore amount of X-ray data collected for imaging), image quality,and/or the precision of manufacturing processes. For example, a FSFOPhaving focused fibers or fiber blocks may be manufactured such that thefocus of the fibers to the radiation source and/or alignment of thefibers to rays originating from the radiation source may be within apredetermined tolerance. As the fibers lengthen and the FSFOP thicknessincreases, the precision of the fiber focus and/or alignment may exceedthe predetermined tolerance. Longer fibers may be more difficult toprecisely focus than relatively shorter fibers. Unfocused or misalignedfibers may reduce image quality, for example, by increasing blur acrossmultiple fibers and array pixels. The FSFOP thickness and/or fiberlength may be selected to balance the need for collecting sufficientdata for dosimetry purposes and for imaging purposes. Furthermore, FSFOPthat exceed a thickness of about 40 cm may encroach on the patient area.

In some variations, the thickness of a fiber optic plate such as a FSFOPmay be selected based on the desired detective quantum efficiency for acertain energy X-ray beam and cost effectiveness. The detective quantumefficiency, DQE(f) is a measure of the combined effects of the signal(related to image contrast) and noise performance of an imaging system,generally expressed as a function of spatial frequency. An ideal imagingsystem would have 100% DQE at all frequencies. For example, a FSFOPhaving 5 cm thickness can provide about 9% DQE(0) at zero frequency fora 6 MV X-ray. For comparison, a conventional EPID, which typically usesCsI(Tl) or Gadolinium oxysulfide (GOS) scintillator screens plus a 1 mmcopper plate, may have about 1% DQE(0). The detector DQE may continue toimprove with increasing fiber plate thickness. For example, at 30 cmfiber plate thickness, the DQE may be 37%. FIG. 1G depicts a plot ofdetection efficiency (which represented the total absorbed X-ray energythat results in a detectable output signal) as a function of thethickness of a detector comprising a water-like material. Higher levelsof detection efficiency may help to facilitate collection of sufficientX-ray data for generating images with higher resolution. FIG. 1Gincludes detection efficiency curves for X-ray or radiation sources ofdiffering energy levels (e.g., 6 MV, 10 MV, 18 MV). For example, a 5 cmthick plate may have about a 18% detection efficiency in a 6MV X-raybeam (which is significantly higher than the 1-2% detection efficiencyof conventional EPIDs that have a copper plate). Increasing thethickness of the fiber optic plate, in combination with the material(s)and/or small diameter of the fibers of the optic plate may promote theconditions that allow a MV detector to function as both a highresolution dynamic imaging detector and as a dosimeter for MV therapyapplications. A FSFOP comprising a water-like material and having athickness of about 5 cm may have a detection efficiency that is suitablefor measuring X-ray data for imaging purposes while also being able tomeasure dose distribution data for a 6 MeV to 25 MeV source.

As described above and in variations throughout this document, thematerial(s) selected for a FSFOP may have water-like properties in orderto measure delivered dose. Other variations of MV detectors may have aFSFOP using fiber material(s) having substantially different propertiesfrom water (e.g., one with higher density, high absorption properties,greater X-ray stopping power, etc.) may not be as suitable. Examples ofsuch materials may include glass, BGO, CWO, CsI, GOS and the like.Because the dose-depth distribution curve of these materials may besubstantially different from that of tissue or water, measuring the dosedistribution in such material(s) may not provide an accurate measurementof dose distribution in tissue. While such fiber material(s) may besuitable for collecting data for generating images (due to a higherX-ray stopping power as compared to water), the precision of the dosedistribution measurements acquired by such MV detectors may becompromised or reduced as compared to MV detectors that use fibermaterial(s) with water-like properties.

FIG. 2 depicts one variation of a MV detector configured to detectradiation for both imaging and dose measurements. The MV detector 200may comprise a fiber optic plate 202, a photodiode array 204 opticallycoupled to the fiber optic plate and configured to receive photons fromthe plate, and data collection/transfer readout electronics incommunication with the photodiode array (not shown). For example, thefiber optic plate 202 may be a FSFOP, the photodiode array 204 may be anamorphous silicon (a-Si) array (e.g., active matrix flat panel imager)or a crystal silicon photodiode and thin-film transistor (TFT) array orother photon imaging devices (CCD, CMOS imaging arrays). The MV detector200 may comprise a left detector module 212 a and a right detectormodule 212 b, each of which may comprise a block or cluster ofscintillating fibers 213 a,b that are focused to a radiation source(e.g., fibers of each block are aligned to different rays of a radiationbeam), and separate photodiode arrays 214 a,b and separate readoutelectronics (not shown). The left and right modules 212 a,b may belocated adjacent to each other to form a rectangular or trapezoidalshape. Although a multi-module MV detector is depicted in FIG. 2 , itshould be understood that an MV detector may not be sub-divided intomultiple modules, and may instead be a single FSFOP unit with a singlephotodiode array and a single unit of readout electronics. In thisexample, the thickness of the fiber optic plate 202 may be about 5 cm,and the length of the MV detector 200 may be about 88 cm.

FIG. 3 depicts another variation of a fiber optic plate that comprisesan array of fiber optic blocks, where each of the fiber optic blocks isfocused to a radiation source. The longitudinal axes of the fibers of afiber optic block may be aligned to a ray of a radiation beam emitted bythe radiation source such that the fiber optic block is focused to theradiation source. In some variations, a fiber optic plate may comprise afirst fiber optic block comprising fibers that are slanted at a firstangle with respect to a top face (e.g., light input face) of the fiberoptic plate such that their longitudinal axes are aligned with a firstray of the radiation source, and a second fiber optic block comprisingfibers that are slanted at a second angle with respect to a top face ofthe fiber optic plate such that their longitudinal axes are aligned witha second ray of the radiation source. The first angle and the secondangle may be different, due to the different locations of the firstfiber optic block and the second fiber optic blocks on the fiber opticplate. The fiber optic plate 300 may comprise a plurality of fiber opticblocks 302 that together form a plate having an overall trapezoidalshape. In the variation of a fiber optic plate of FIG. 3 , the fiberoptic plate 300 comprises 48 (3×16) blocks of focused scintillatingoptical fibers, where each of the 16 blocks are focused to different orunique rays 305 a, 305 b, 305 c of a radiation beam 305 emitted from theradiation source 304. Although the fiber optic plate 300 is depicted ashaving fiber optic blocks 302 that are focused to the radiation source(which may be represented by a focal spot) 304 in two dimensions (e.g.,along the x and z directions), it should be understood that the fiberoptic blocks 302 may be focused to the radiation source 304 in threedimensions (e.g., along the x, y and z directions). The radiation sourcemay be any distance above the plate, and the patient area may be locatedin the space between the radiation source 304 and the fiber optic plate300. For example, the distance D between the radiation source and the MVdetector (e.g., the fiber optic plate) may be from about 110 cm to about150 cm, e.g., about 132.5 cm. As an example, fiber optic plate 300 mayhave a thickness T of about 5 cm thick, a length L of about 88 cm, and awidth W of about 8 cm.

The individual optic fibers and/or fiber optic blocks of a fiber opticplate may be directly coupled to (e.g., contacting) each other, withoutany intermediary light-shielding between the fibers and/or the blocks.Optionally, the fiber optic plate may comprise light-shielding layers orsepta or cladding between individual fibers and/or between the fiberoptic blocks. The light-shielding layer or coating may help to reducethe scatter of X-rays and/or visible light photons from one block and/orfiber to another adjacent block and/or fiber. In some variations, alight-shielding layer that is interposed between blocks and/or fibersmay comprise thin sheets of high-Z metals, for example, 0.1 mm thicktungsten or lead. The thickness of the septa or cladding (e.g., thedistance between each block) can be made smaller or larger to trade-offbetween scatter effectiveness and the loss of incoming primary X-raysdue to the space occupied by the septa or cladding. For example, thethickness of the septa or cladding may be selected such that the noisearising from inter-fiber scatter and the image blur due to the loss ofprimary X-rays are able to be corrected using image processing methods.In some variations, a fiber optic plate may comprise any light-shieldinglayers or septa or cladding between the individual fibers and/or betweenthe fiber optic blocks. The scatter of X-ray and/or visible lightphotons between fibers may be reduced or blocks by using fibers orwaveguides made of materials having different refractive indices andarranging them in an alternating or randomized fashion in the plate sothat fibers with different refractive indices are adjacent to each other(e.g., a Tranloc waveguide). The interface between two materials havingdifferent refractive indices tends to cause light to reflect at theinterface, instead of being transmitted through from one material to theother. By arranging fibers having different refractive indices adjacentto each other, light in one fiber is kept from crossing over to anadjacent fiber. In such matter, the interface between two fibers withdifferent refractive indices may act as a light barrier and may help toreduce light scatter between fibers. In one variation, a fiber opticplate (e.g., FSFOP) may be comprise fibers made of a first material witha first refractive index (e.g., polystyrene) and fibers made of a secondmaterial with a second refractive index (e.g., acrylic or poly methylmethacrylate). The fiber optic plate may be made by drawing these twofiber types together such that the fibers are arranged in an alternatingor semi-randomized or randomized fashion. A fiber optic plate comprisingfibers or waveguides made of two or more materials with differentrefractive indices (e.g., a Transloc waveguide) may reduce or eliminatethe use of cladding or septa between the fibers or blocks, which mayreduce the amount of “dead space” occupied by cladding or septa on thesurface of the light sensor array.

Other components may optionally be included in a fiber optic plate tohelp reduce the incidence of scattered X-rays on the photodiode arrayand/or shield the MV detector from external noise sources, such as fromthe linac, 60 Hz noise from power supplies, and the like. For example, aMV detector may comprise a fiber optic plate and a cover located on thetop surface of the fiber optic plate (i.e., the surface of the MVdetector that is closest to the radiation source). The cover may be athin metal plate made of low-Z metal, for example, 0.2 mm Cu or 1 mm Al.The cover may be electrically grounded, and may be configured to blockmore scattered X-rays than primary X-rays. Alternatively oradditionally, a reflective paint material or coating may be appliedbetween the individual fibers and/or between blocks, which may help toreflect scintillation photons back into the fibers and toward thephotodiode array. Alternatively or additionally, a reflective paintmaterial may be applied on the fiber plate surface opposite to thephotodiode/TFT array to reflect the light photons towards thephotodiode/TFT array.

As described above, the fiber optic plate may contact and interface withthe light sensor array (e.g., photodiode/TFT array) such that visiblelight channeled through the fibers (or in some variations, generated bythe fibers themselves) impinges on the light sensor array, whichconverts the photon incidence to an electrical signal. In somevariations, as described above, the fibers of the fiber optic plate maybe much smaller than the pixels of the photodiode/TFT array such thatmultiple fibers map onto (e.g., contact) a single pixel. FIG. 4 depictsone example of how a fiber optic plate may map onto a photodiode array.The thickened dotted lines 402 represent the pixel boundaries on aphotodiode array. The boundaries of the individual fibers of the fiberoptic plate (e.g., a plastic scintillator array) are represented by thethin lines 404. The fibers 404 have a much smaller diameter than thepixel width of the photodiode/TFT array (e.g., where the ratio betweenthe fiber diameter to the pixel width is about 1:20). Instead of aone-to-one mapping between the fibers and the pixels (i.e., the fibersof a fiber optic plate and the pixel boundaries of a photodiode arrayare not aligned such that one each pixel is in contact with exactly onefiber), multiple fibers map onto a single pixel and the uniformity ofeach pixel may be achieved by the averaging effect. That is, lightsignals from multiple fibers impinging on a single photodiode/TFT pixelmay be summed and averaged by that photodiode/TFT. In this example, thediameter of the fibers may be from about 5 μm to about 50 μm (e.g.,about 5 μm to about 10 μm, about 10 μm), while the photodiode arraypitch or pixel width may be from about 150 μm to about 600 μm (e.g.,about 400 μm). In some variations, a single pixel may be in contact withand/or receive light data from anywhere between 10 to 200 fibers, e.g.,10, 20, 25, 30, 50, 100, 10-100, 50-200, etc., fibers. The fiber opticplate may be coupled to the light sensor array such that there islittle, if any, air at the plate-array interface. For example, one ormore clamps around the perimeter edge of the fiber optic plate may beused to secure the fiber optic plate to the light sensor array.Optionally, the surfaces of the fiber optic plate and the light sensorarray that contact each other may be polished to help facilitate anair-free interface between the fiber optic plate and sensor array.Alternatively or additionally, adhesives may be provided between thefiber optic plate and the light sensor array (e.g., around theperimeter) to secure these components together.

The light data collected by the photodiode/TFT array (which may be ana-Si detector) may be in communication with readout electroniccircuitry, which are schematically depicted in FIG. 5 . In somevariations, an a-Si photodiode/TFT array or a-Si detector 500 may be incommunication with electronic circuitry 502 configured for dynamichigh-frame-rate readout. The a-Si detector 500 may comprise a firstdiode array 501 (e.g., left diode array) and a second diode array 503(e.g., right diode array). The readout electronic circuitry 502 maycomprise one or more PCB or electronic substrates 507, and a pluralityof data connectors 504 configured to capture a dose distribution imagefor each linac beam pulse (which may be about 100-200 frames per second)and to transmit detector data to electronic components (e.g.,machine-readable memories, multiplexors, microprocessors, and the like)that may be mounted on the substrate 507. In one variation, readoutelectronic circuitry may comprise additional connectors 506 for powerand/or control signals. Connectors 506 may include connectors forsignals that contain timing synchronization data (e.g., triggersignals), connectors for network interface and data exchange (e.g.,Ethernet connections, GigE-capable connections), and/or connectors forpower. The system controller and/or linac may provide a sync signal tothe readout electronic circuitry to coordinate the timing between theapplication of the radiation beam and the acquisition of light data fromthe MV detector. The interconnect between the a-Si photodiode/TFT array500 and the readout electronic circuitry 502 may comprise a flexiblewired bus or ribbon or tape 508 having conducting traces (e.g., copper,gold, etc.) embedded in an insulating material, which may be bonded tothe detector array and/or connectors using solder or adhesives (e.g.,anisotropic conductive adhesives such as anisotropic conductive film orpaste). In some variations, the length 509 of the bus 508 between thediode array 501 and the substrate 507 may be about 4 cm and the width511 of the substrate 507 may be about 10 cm. The a-Si photodiode/TFTarray may have a native pixel resolution of about 100 μm to about 200μm, and may have an overall length 510 of about 85 cm, e.g., about 86.73cm and a width 512 of about 10 cm. The pixels may be binned to 2×2, 3×3or 4×4 matrices during the readout to speed up the frame rate and toreduce the image data size. In some variations, the whole photodiodearray may consist of a matrix of 4400×400 pixels at the pixel pitch of0.2 mm. During or after the readout, the pixels can be binned in variousformats. For example, with a 2×2 binning format, the resulting newmatrix would have 2200×200 pixels at pixel pitch of 0.4 mm. Binning thepixels may result in proportionally larger effective pixel pitch, orlower resolution.

The controller (e.g., the radiation therapy system controller) mayprovide control/sync signals to the X-ray source (e.g., linac) and MVdetector to synchronize the operation of these components. The controlsignals may also indicate the mode of operation of the system, and theacquired image data may be sorted, stored, and processed in accordancewith the operating mode. One example of a MV detector and linac timingdiagram 600 is illustrated in FIG. 6 . As depicted there, a linacmodulator trigger signal 602 may be generated by the radiation therapysystem controller, which may be transmitted to the linac and the MVdetector so that the radiation data collection is synchronized with theradiation beam firing event. In this variation, the linac modulatortrigger signal 602 is a constant periodic signal that triggers the MVdetector to acquire an image at a set time interval, regardless ofwhether the linac has fired a radiation beam or not, but in othervariations, the trigger signal 602 may not be synchronous or periodic.The read sync signal 608 synchronizes the data output from thephotodiode/TFT array with the data connectors 504. In some variations,the read sync signal 608 may be activated in accordance with the triggersignal 602 (e.g., after a delay), and may represent the transfer ofimage data from the detector to the controller, which may be offset by adelay after the trigger signal. Image data may be raw data from thephotodiode/TFT array, or may be the average of several acquired images(one averaging method is described below and depicted in FIGS. 12A-12B).When the controller transitions to the treatment mode from thenon-treatment mode (as indicated by the treatment signal 604transitioning from a low value to a high value, at time point t0, forexample), the controller may send a linac beam pulse 606 to the linac tofire a radiation beam (e.g., X-ray). In this variation, as long as thetreatment signal 604 has a high value (indicating that the system is inthe treatment mode), the linac beam pulse 606 will pulse according tothe trigger signal 602. While the controller is in treatment mode (asindicated by duration 609), the MV detector acquires data that may beused to generate an image or plot of the radiation deposited in thepatient. When the controller is not in treatment mode, the datacollected by the MV detector may be used to generate “dark” or “offset”images. Dark or offset images may be used to correct background or noiseartifacts. For example, the data and/or images acquired in the absenceof a linac radiation beam may be subtracted out from the data and/orimages acquired in the presence of a linac radiation beam to compensatefor the effects of detector dark current.

Described herein are methods of manufacturing fiber optic plates for aMV detector that may be used for both anatomical imaging and radiationdose measurements. As described above, a fiber optic plate may comprisea plurality (e.g., an array) of optical fiber blocks or bundles coupledtogether. The individual blocks may be manufactured separately and thenassembled together to form a complete optical fiber plate. For example,a FSFOP may comprise a plurality of scintillating fiber optic blocks,where each block comprises fibers having longitudinal axes that arealigned to a ray of a radiation beam emanating from a radiation source.Different blocks may comprise fibers with longitudinal axes that arealigned to different rays of the radiation beam, such that each of theblocks is focused to the radiation source. The size and shape of theFSFOP may vary to correspond with the size and shape of the radiationbeam and/or the physical arrangement of a radiation therapy or imagingsystem. Accordingly, the number and size of the fiber optic blocks mayalso vary, and each block may have a width from about 1 cm to about 10cm. Additionally, the angle of the fibers in a fiber optic block mayvary depending on the radiation ray to which the block is focused andthe location of the block in the overall fiber optic plate. In somevariations, at least one fiber in a fiber optic block 700, e.g., thelongitudinal axis of the central fiber as depicted in FIG. 7A, isfocused to the X-ray source. One method of manufacturing a fiber opticblock 700 that has at least one focused fiber, such as the blockdepicted in FIG. 7A, is depicted in FIGS. 7B-7C. Each block of a fiberoptic plate may be individually cut from a plate of parallel fibers at adifferent specific tilt angles such that the central fibers of each ofthe blocks are focused. The tilt angle at which each of the blocks arecut may depend on the location of the particular block in the fiberoptic plate of the detector assembly. In some variations, only thefiber(s) located near or at the center of the block are focused. Thefibers that are further from the center of the block become less andless focused (e.g., deviate more and more from the rays of a radiationsource) to the radiation source. In one variation of a method 720 formanufacturing a block or module of fiber optic plate comprising aplurality of blocks or modules, a grouping or bundle of fibers that aresubstantially parallel may be manufactured or procured 722, according toknown methods. The fiber tilt angle (e.g., the angle of the fibers inthe block with respect to the line that is normal to the top and bottomsurfaces of the block) of each of the fibers in the block may bedetermined at least in part based on the radiation ray to which theblock is to be focused and/or the intended location of the block in thefiber optic plate, and may, for example, be from about 0 degrees toabout 90 degrees, e.g., about 0 to about 60 degrees. The method 720 maycomprise creating 724 a top surface 705 by slicing along a first line704 that transects across the fibers at an angle A1 (e.g., the tile facecut angle), creating 726 a bottom surface 707 by slicing along a secondline 706 that transects across the fibers at an angle A2 and is parallelto the top surface 705 (i.e., angles A1 and A2 are substantially thesame). The second line 706 may be a distance D1 away from the topsurface 705. The method 720 may also comprise creating 728 a first sidesurface by slicing along a line 710 a, which is at a tile cut angle A3(which is the angle from a line that is normal to the top surface of theblock), creating 730 a second side surface by slicing along a line 710b, which is at a tile cut angle A4, creating 732 a third side surface byslicing along a line 710 c which transects the top and bottom surfacesat a tile cut angle A5, and creating 734 a fourth side surface byslicing along a line 710 d which transects the top and bottom surfacesat a tile cut angle A6 (see FIG. 7A). The tilt cut angles A3-A6 may varyfor each of the side surfaces, depending on the location of theparticular optical fiber block within the entire fiber optic plate, andthe overall shape (e.g., trapezoidal or cuboid shape) of the fiber opticplate. These steps may be repeated for each block of the plurality ofblocks of the fiber optic plate, but the tile cut angles and tilt cutangles may vary for each of the blocks. In the example depicted in FIG.7A, at least the longitudinal axis of central fiber 709 is aligned toradiation ray 712, while the longitudinal axes of the fibers adjacent tothe central fiber 709 may not be exactly aligned to ray 712. The numberor proportion of fibers in a particular block of a fiber optic plate mayvary, and may be selected such that the spatial distortions of theresultant images are in the sub-millimeter range (e.g., no more thanabout 1 mm). In some variations, all (about 100%) of the fibers may befocused, while in other variations, the ratio of focused fibers tounfocused fibers in a block may range from all the fibers being focused,to most of the fibers being focused (e.g., from about 100:1 to about2:1), to about half of the fibers being focused (e.g., about 1:1). Thefocused fibers may be distributed across the block, or may beconcentrated in the central portion of the block.

In some variations, a fiber optic plate may comprise a plurality offiber blocks, where most, if not all, of the fibers in each of theblocks is focused to a radiation source (e.g., where the longitudinalaxis of each fiber is aligned along a ray of the radiation beam). Forexample, all of the fibers in a block may be focused (e.g., 100% of thefibers) or the proportion of focused fibers to unfocused fibers may befrom about 100:1 to about 2:1 (e.g., about 99% to about 51% of thefibers are focused). Such blocks may be considered “fully focused”blocks, since most or all of the fibers are focused to the radiationsource. FIGS. 8A and 8B depict one variation of a method 820 formanufacturing a fully focused fiber optic block 800. The method 820 maycomprise procuring or manufacturing a tapered fiber billet. The fiberbillet may be thermally tapered, for example. The geometry for taperingfollows the focal distance and location of the each fiber in the fiberblock.

FIG. 8C shows the geometry for one fiber, where the center of the topfiber plate surface is at the origin, x and z are the distances of thefiber from the origin, and y is the radiation source (e.g., X-ray focalspot) distance from the origin. The tilt angle of each fiber, θ, isdependent of its (x, z) coordinates and is calculated by the followingequation,

$\theta = {\arctan\left( \frac{\sqrt{x^{2} + z^{2}}}{y} \right)}$

The method 820 may comprise creating 824 a top surface 805 by slicingalong a first line 804 that transects across the fibers at an angle A7(e.g., the tile face cut angle), creating 826 a bottom surface 807 byslicing along a second line 806 that transects across the fibers at anangle A8 and is parallel to the top surface 805 (i.e., angles A7 and A8are substantially the same). Angles A7 and A8 may be selected in orderfor one or more fibers in the block to have a tilt angle A13 inaccordance with Equation 1. The second line 806 may be a distance D2away from the top surface 805. The method 820 may also comprise creating828 a first side surface by slicing along a line 810 a, which is at atile cut angle A9 (which is the angle from a line that is normal to thetop surface of the block), creating 830 a second side surface by slicingalong a line 810 b, which is at a tile cut angle A10, creating 832 athird side surface by slicing along a line 810 c which transects the topand bottom surfaces at a tile cut angle A11, and creating 834 a fourthside surface by slicing along a line 810 d which transects the top andbottom surfaces at a tile cut angle A12 (see FIG. 8D). The tilt cutangles A9-A12 may vary for each of the sides, depending on the locationof the particular optical fiber block within the entire fiber opticplate, and the overall shape (e.g., trapezoidal or cuboid shape) of thefiber optic plate. In some variations, the tilt angle for any fiberwithin the fiber block may be determined at least in part by Equation 1.These steps may be repeated for each block of the plurality of blocks ofthe fiber optic plate, but the tile cut angles and tilt cut angles mayvary for each of the blocks.

FIGS. 9A-9B and 10A-10B depict one example of a plurality of fiber opticblocks 1002 assembled together to form a fiber optic plate 1000 that maybe focused to a radiation source 1001, which may be represented as afocal spot. FIG. 9A depicts one portion (e.g., the right half) of thefiber optic plate 1000 comprising assembled fiber blocks 1002 along theX-direction, FIG. 9C depicts a table (Table 1) that outlines exampledimensions and angles of the portion of the fiber optic plate 1000depicted in FIG. 9B (which is a reproduction of FIG. 9A for the purposesof clarity). The fiber optic plate 1000 may optionally comprise a low-Zlayer or coating disposed over the top surface of the plate (e.g., thesurface that faces the radiation source 1001), which may help to deflectlow-energy, scattered X-rays, as described above. FIG. 10A depicts thefiber optic plate 1000 along the Z-direction and FIG. 10B depicts atable (Table 2) that outlines example dimensions and angles of theportion of the fiber optic plate 1000 depicted in FIG. 10A. The overallgeometry of the fiber optic plate 1000 may be similar to that of theplate depicted in FIG. 1 . This particular fiber optic plate 1000 maycomprise eight blocks 1002 (numbered B1-B8) along the X-direction, withor without a thin sheet of heavy metal, which may be used asanti-scatter septa. Thin sheets of heavy metal may be located at theinterface 1004 between each block 1002. Optionally, a thin sheet ofmetal may also be located along at least a portion of the top surface1006 of the plate. The dimensions of each block may vary depending onthe overall size of the fiber optic plate 1000. For example, the widthW1 of each block B1-B8 along its bottom edge may be about 5.5 cm for anoverall width in the x-direction of about 88 cm (about 44 cm for theleft and right half). An MV detector with an overall width of about 88cm may be used to detect the dose distribution and/or measure imagingdata for a 6 MV radiation source located about 130 cm away from thedetector, e.g., about 132.5 cm. The distance 1021 to the peripheral edgeof each of the blocks B1-B8 from the central edge 1020 is summarized inTable 1. The distance 1023 to the block center from the central edge1020 is also summarized in Table 1. The width W2 of each block along itstop edge may be about 2.67 cm for an overall width in the z-direction ofabout 8 cm. The thickness T1 of the fiber optic plate may be about 5 cm.The left side blocks are symmetric and identical to the right side. Thecenter block 1008 may have a fiber tilt angle 1010 of about 0 degreesand the tilt angle 1011 a-1011 g may progressively increase towards theend block 1012, where the fiber tilt angle in this angle is representedby the angle of the focused fibers of each fiber block relative to thetop surface 1006. Examples of specific fiber tilt angles are alsosummarized in Table 1 of FIG. 9C. The blocks 1002 may be made fromeither the method depicted in FIGS. 7A-7C or the method depicted inFIGS. 8A-8D. The central fiber of each block may be focused to aradiation source, and/or the block may be fully focused, as previouslydescribed.

Radiation therapy systems that comprise the MV detectors described abovemay have at least two modes of operation. These modes of operation maybe used in pre-treatment methods and/or may be used during and/or aftertreatment. A first mode of operation may be a “step-and-shoot” mode,where the gantry upon which the MV detector and radiation source aremounted is stationary during the imaging (e.g., not moving) and a singlehigh dose projection image can be taken at a selected angle. The gantrycan rotate to multiple angles and projection images can be taken fromeach angle. A second mode of operation may be a tomo-graphic mode, wherecontinuous low-dose, thin slices of image are taken while the gantry isrotating, i.e. the gantry is rotating while the MV detector is acquiringimages. Both of these modes may be used for in a pre-treatment methodfor patient positioning. For pre-treatment imaging, the radiation sourcemay emit a lower-energy beam (e.g., about 3 MV), which may result in animage with better contrast than using a high-energy beam (e.g., about 6MV) which is used during treatment. While the MV detectors describedherein may be used for dose measurements, in some systems, the data fromthe MV detectors may be used for imaging purposes, and may notnecessarily be used for generating dose maps or images.

During treatment, the MV detector may continuously measure the radiationthat is transmitted through the patient (e.g., according to the timingdepicted in FIG. 6 ). Based on the radiation applied by the radiationsource and the radiation transmitted through the patient, the systemcontroller may be configured to determine the radiation dose deliveredor absorbed by the patient. In some variations, a method for radiationtherapy may comprise moving the patient couch slowly through the patientregion of the gantry while the gantry is rotating so that the radiationsource emits radiation beams from multiple angles. The patient couch maymove at very slow speed, for example about 0.07 mm/second. At this rateof movement, the helical pitch (couch-displacement perrotation/beam-width) is much smaller than 1 mm. Table 3 depicted in FIG.11A provides some example parameters (e.g., rates of rotation andmovement) for one variation of a treatment session. Additional detailsregarding the computation of delivered dose are provided below anddepicted in FIGS. 11B-C.

When the movement of the patient couch through the fan beam issignificantly slower than the rate of gantry rotation, large portions ofthe slice may overlap each other. A slice is defined as the couch traveldistance per rotation of the gantry. During the treatment session, theMV detector collects measurements of the radiation dose applied to thepatient in-vivo and in real-time. That is, the MV detector measures theamount of radiation applied by the radiation source at each firingposition around the gantry. For each firing position, the overlappedportion of the images can be summed and averaged. An anatomical regionof interest may be imaged from the same angle multiple times during theslow couch motion. In each image, the anatomical region of interest maymove at the same rate that the couch moves. These images may be shiftedto overlap with each other. The overlapped images can then be summed andaveraged. Such an image averaging scheme may help to improvesignal-to-noise ratio (SNR). The averaging process can be carried outwith an on-board computer, thereby reducing the amount of image datathat needs to be transmitted to a remote computer and archived.

One example of an image averaging scheme is illustrated in FIGS. 12A-12Band 13 . In this example, the beam width is 8 mm, the lateral tablemovement per rotation (Z) is 0.07 mm/rotation, so the accumulatedsame-angle images are 8/0.07=114 rotations. Therefore 114 images can beaveraged at any fixed angle of rotation. In this example, the SNR mayimprove by factor of 10.6, and image data volume may be reduced byfactor of 114. In the example depicted in FIG. 12A, the slice width maybe about 1 cm and couch displacement may be about 0.0704 mm per gantryrotation. During 1 cm of couch movement the gantry completes 142rotations and for each rotation the MV detector acquires 100 projectionimages from the 100 firing angles. The MV detector may also acquire 142view of the same anatomical slice, but shifted across the detectorplane. FIG. 12B is a representation of a wave-shaped profile ofprojection images shifted across the detector plane. The images from thesame gantry angle may be spatially shifted and overlapped with oneother. The overlapped images may be averaged to improve the signal tonoise ratios of the images. FIG. 13 is a graphical representation of thenumber of images available for averaging as a function of couchdisplacement for one variation of a radiation therapy system. The shadedareas show the start and end fringe regions. For first and the lastcentimeter (1 cm) of the couch movement, the number of images availablefor averaging will be less than 142 images. This is because the sameangle views have not accumulated to maximum of 142.

Because the detector is made of water equivalent plastic materials, thedosimetric properties of the detector may facilitate dose calculationwithout the need for computationally-intensive corrections, as comparedto the existing non-water equivalent EPID detectors. After a treatmentsession, a projection image may be generated from the dose calculationand combined with an anatomical image of the patient to generate a 3-Dpatient dose distribution image. In some variations, a controller of aradiation therapy system with any of the MV detectors described hereinmay be configured to generate a dose distribution map or image after atreatment session. For example, the radiation therapy system may beconfigured to acquire an anatomical image before the treatment session,acquire radiation dose data during the treatment session, process theradiation dose data (e.g., during or after the treatment session), andgenerate a composite image or map after the treatment session thatcomprises an overlay of the radiation dose distribution map or imageover the pre-treatment anatomical image. Optionally, the radiationtherapy system may acquire an anatomical image immediately after thetreatment session and generate a composite image or map that comprisesan overlay of the radiation dose distribution map or image over thepost-treatment anatomical image. In some variations, the pre-treatmentanatomical image and post-treatment anatomical image may be averagedtogether for the composite anatomical-dose distribution image.Alternatively, the radiation therapy system may be configured togenerate a radiation dose distribution map or image, and not generate ananatomical image. Any anatomical images acquired by the radiationtherapy system may be acquired using either (or both) 3 MV or 6 MVX-rays. For example, anatomical images acquired before or after thetreatment session may be acquired using 3 MV X-rays, while anyanatomical images acquired during the treatment session (e.g., whenradiation beams for treatment are being applied to the patient) may beacquired using 6 MV X-rays.

One variation of a method for computing a dose map is depicted in FIG.11C. The method 1100 may be used with a radiation therapy system that isconfigured to generate narrow beams (e.g., pencil beams) that are on theorder of about 1 cm². Such narrow beams may be generated, for example,by collimating the output radiation of an X-ray source with a binarymulti-leaf collimator (MLC). In some variations, the radiation therapysystem may also comprise a rotatable gantry upon which the X-ray orradiation source is mounted and an MV detector (such as any of thevariations described herein) disposed across from the X-ray source, suchas the system depicted in FIG. 1A. The gantry may be rotatable about apatient area. Data from the MV detector may be stored and processed in asystem controller comprising a microprocessor and computer-readablememory. In some variations, the steps of the method 1100 may be storedin a computer-readable memory as a pre-determined instruction set. Inaddition, the rotation of the gantry, firing of the X-ray or radiationsource, MLC control, data output from the MV detector, including thetiming related to these steps (e.g., firing and collimating radiation atcertain gantry angles), may be carried out according to control signalsoriginating from the system controller. The method 1100 may comprise thestep 1102 of acquiring images of radiation beams in the absence of apatient (or a phantom) to generate entrance beam image(s). FIG. 11Bschematically depicts an entrance beam 1130 as the radiation emitted bythe X-ray or radiation source 1131 before it interacts with a patient ora phantom (which may be located in the region 1132 enclosed in dottedlines; in this particular depiction, there is no patient or phantom).Entrance beam images may include images of all the MLC single andmultiple adjacent leaves simultaneously opening and may be acquired froma plurality of gantry or firing angles. The entrance beam image(s) maybe acquired at one firing angle or at multiple firing angles. In somevariations, the number of images may be reduced by limiting the numberof adjacent leaves that can be opened during treatment (e.g., up to 6adjacent leaves opening). Entrance beam image(s) may represent the X-rayfluence without any beam attenuation for each MLC leaf. The radiationincident upon the plane of the MV detector may be scaled to a patiententrance surface using, for example, the inverse radius square law.

After entrance beam images are generated, method 1100 may comprise thestep 1104 of acquiring images of radiation beams in the presence of apatient (and/or a phantom) at a first couch position to generate exitbeam image(s). Step 1104 may comprise loading a patient or phantom on acouch and advancing the couch into the patient area of the gantry. FIG.11B schematically depicts an exit beam 1134 as the radiation thatemerges from the patient (or phantom) that is incident on an MV detector1136; i.e., radiation that has interacted with the patient or phantom.Step 1104 may be performed, for example, during a therapy session, andmay include acquiring exit beam images of all the MLC single andmultiple adjacent leaves simultaneously opening. The exit beam image(s)may be acquired at one firing angle or at multiple firing angles thatcorrespond to the angles acquired in step 1102. In some variations, exitbeam images may be averaged in order to reduce the number of images (forexample, using the method depicted in FIGS. 12A-12B). The images of exitX-ray fluence from the patient may be recorded in vivo, i.e., in realtime, by the MV detector for all the firing positions around the gantry.The exit beam images may be generated during couch movement, forexample, as the couch moves in accordance with the parameters summarizedin Table 3 of FIG. 11A. For example, the couch may move about 100 mm ata speed of about 0.07 mm/s. As explained previously, the total number ofimages (142857) may be reduced to 1000 by averaging the overlappedimages. Alternatively or additionally, step 1104 may be executed whilethe couch is stopped at the first couch position.

After exit beam images are generated, method 1100 may comprise a step1106 of computing attenuated fluence image(s) based on the entrance beamimage(s) and exit beam image(s). In some variations, attenuated fluenceimage(s) (e.g., patient attenuated fluence images) may be generated bysubtracting the averaged exit beam images of each corresponding leavesfrom the entrance beam leaf images.

Method 1100 may comprise the step 1108 of converting the attenuatedfluence image(s) to attenuated dose image(s). The attenuated fluenceimages may be converted to attenuated dose images by empiricalcalibrations. One example of a calibration method may comprise placingdosimetric films (which may provide 2D dose images) in the same positionof the MV detector imaging photodiode array plane. The film may beplaced inside water or plastic phantoms with the same or equivalentthickness as the plastic FSFOP layer and the same back scatter layer.The film dose image and the MV detector fluence image may be acquiredunder the identical MV beam conditions (e.g., beam energy, attenuationlayer, exposure time, etc.). The calibration method may further comprisecreating correlation maps between the 2D MV detector fluence images andthe corresponding the 2D film dose images. Because the MV detectorsdisclosed herein use low-Z, water-like scintillating materials, thisconversion may be done without the complex calibration procedures usedin for detectors (such as EPIDs) that comprise high-Z scintillationmaterials. Optionally, method 1100 may comprise the step 1110 offiltering out radiation noise (e.g., scatter correction). ScatteredX-rays from the patient may create noise in the images and may causeerrors in dose calculations. The amount of scatter may be proportionalto the area irradiation by the radiation beam after it has been shapedby a collimator. In case of single leaf beam geometry (e.g., the beam isshaped by a multi-leaf collimator), the beam size may be less than fewsquare centimeters (e.g., about 1 cm×0.625 cm or about 2 cm×0.625 cm atISO), the scatter may be negligible and the scatter corrections may notbe needed. For multiple leaf openings, scatter correction may help toimprove the accuracy of the dose calculations. A scatter correctionmodel may be created based on Monte Carlo simulations, or based onempirical scatter measurements, or combination of both.

Method 1100 may comprise the step of 1112 of determine the dosedistribution along each firing angle (e.g., firing position) at thefirst couch position. For example, some systems may have about 100firing positions around a gantry (e.g., a circular gantry), where eachfiring positions are about 36 degrees from each other. In somevariations, the dose distribution may be computed by back projecting theexit dose image to patient dose along each firing angle for all openedleaves. The dose distribution along each firing angle may be computedusing the preexisting data sets, which include PDD curves, beamgeometry, energy spread kernel in the patient, and may optionallyinclude data from patient CT images of attenuation coefficient numbers,i.e., CT numbers. A 2D patient dose may be calculated by the process ofconvolution and super position from all the firing angles in singleslice. Notably, in this dose reconstruction method, the correction orcalibration steps needed for the EPID dose images are not necessary,since the MV detectors here use low-Z scintillating fibers.

Method 1100 may comprise the step 1114 of moving the patient couch to asecond position and repeating steps 1104-1112 at the second position(i.e., the second slice). Steps 1104-1112 may be repeated 1116 for alldesired patient and couch positions. Method 1100 may comprise the step1118 of computing a cumulative dose distribution map or image (which maybe 3D) based on the dose distribution computations from steps 1104-1116for all desired patient and couch positions. One of more of the steps ofthe method 1100 may be performed during a treatment session (e.g., whilea patient is on the couch) and/or after a treatment session (e.g., aftera patient has left the couch).

Although the foregoing systems, devices and methods have, for thepurpose of clarity and understanding, been described in some detail byway of illustration and example, it will be apparent that certainchanges and modifications may be practiced, and are intended to fallwithin the scope of the appended claims.

Where a range of values is provided, it is understood that eachintervening value, to the tenth of the unit of the lower limit unlessthe context clearly dictates otherwise, between the upper and lowerlimits of that range is also specifically disclosed. Each smaller rangebetween any stated value or intervening value in a stated range and anyother stated or intervening value in that stated range is encompassedwithin the invention. The upper and lower limits of these smaller rangesmay independently be included or excluded in the range, and each rangewhere either, neither or both limits are included in the smaller rangesis also encompassed within the invention, subject to any specificallyexcluded limit in the stated range. Where the stated range includes oneor both of the limits, ranges excluding either or both of those includedlimits are also included in the invention.

Unless defined otherwise, all technical and scientific terms used hereinhave the same meaning as commonly understood by one of ordinary skill inthe art to which this invention belongs. Although any methods andmaterials similar or equivalent to those described herein can be used inthe practice or testing of the present invention, some potential andpreferred methods and materials are now described.

It must be noted that as used herein and in the appended claims, thesingular forms “a”, “an”, and “the” include plural referents unless thecontext clearly dictates otherwise. Thus, for example, reference to “ablock” includes a plurality of such blocks and reference to “the pixel”includes reference to one or more pixels and equivalents thereof knownto those skilled in the art, and so forth.

The publications discussed herein are provided solely for theirdisclosure. Nothing herein is to be construed as an admission that thepresent invention is not entitled to antedate such publication by virtueof prior invention. Further, the dates of publication provided, if any,may be different from the actual publication dates which may need to beindependently confirmed.

The preceding merely illustrates the principles of the invention. Itwill be appreciated that those skilled in the art will be able to devisevarious arrangements which, although not explicitly described or shownherein, embody the principles of the invention and are included withinits spirit and scope. Furthermore, all examples and conditional languagerecited herein are principally intended to aid the reader inunderstanding the principles of the invention and the conceptscontributed by the inventors to furthering the art, and are to beconstrued as being without limitation to such specifically recitedexamples and conditions. Moreover, all statements herein recitingprinciples, aspects, and embodiments of the invention as well asspecific examples thereof, are intended to encompass both structural andfunctional equivalents thereof. Additionally, it is intended that suchequivalents include both currently known equivalents and equivalentsdeveloped in the future, i.e., any elements developed that perform thesame function, regardless of structure. The scope of the presentinvention, therefore, is not intended to be limited to the exemplaryembodiments shown and described herein. Rather, the scope and spirit ofpresent invention is embodied by the appended claims. For all theembodiments described herein, the steps of the method need not beperformed sequentially.

1-53. (canceled)
 54. A method for computing a dose map, the methodcomprising: acquiring images of radiation beams from a radiation sourcein the absence of a patient on a platform to generate entrance beamimages; acquiring images of radiation beams from the radiation source inthe presence of the patient on the platform to generate exit beamimages; computing attenuated fluence images based on the entrance beamimages and the exit beam images; converting the attenuated fluenceimages to attenuated dose images; determining a dose distribution alongeach firing angle of the radiation source for each of a plurality ofplatform positions using the attenuated dose images; and computing acumulative dose distribution map based on the dose distribution for eachfiring angle and plurality of platform positions.
 55. The method ofclaim 54, further comprising filtering the attenuated fluence images toreduce radiation noise before converting the attenuated fluence imagesto the attenuated dose images.
 56. The method of claim 54, whereinacquiring images of radiation beams comprises emitting radiation fromthe radiation source and detecting the emitted radiation using aradiation detector comprising a fiber optic array having a plurality ofscintillating fibers, wherein the scintillating fibers are focused tothe radiation source.
 57. The method of claim 54, wherein each of thescintillating fibers has a diameter, an input face, an output face, anda longitudinal axis therebetween, and wherein the radiation detectorfurther comprises a photodiode array coupled to the output faces of thefibers in the array, the photodiode array comprising a plurality ofpixels each having a pixel width, and wherein the fiber diameter issmaller than the pixel width, and a plurality of output faces of aplurality of fibers contact each pixel of the photodiode array.
 58. Themethod of claim 54, wherein acquiring images of radiation beamscomprises emitting radiation from the radiation source and detecting theemitted radiation using a radiation detector comprising an optical fiberplate comprising a plurality of optical fiber modules having a pluralityof scintillating optical fibers, where a central scintillating opticalfiber of each of the optical fiber modules is focused to the radiationsource.
 59. The method of claim 58, wherein a longitudinal axis of thecentral scintillating optical fiber of each optical fiber module isangled toward the radiation source.
 60. The method of claim 54, whereinthe entrance beam image and the exit beam image each comprise radiationdetector data acquired from one or more firing angles.
 61. The method ofclaim 60, wherein a multi-leaf collimator comprising a plurality ofleaves is disposed in front of the radiation source, and the entrancebeam images and the exit beam images comprise radiation detector dataacquired with a single open leaf.
 62. The method of claim 61, whereinthe entrance beam images and the exit beam images comprise radiationdetector data acquired with multiple open adjacent leaves.
 63. Themethod of claim 61, wherein computing the attenuated fluence imagescomprises subtracting exit beam images from entrance beam images of thesame leaf openings.
 64. The method of claim 61, wherein computing theattenuated fluence images comprises averaging exit beam images,averaging entrance beam images, and subtracting the averaged exit beamimages from the averaged entrance beam images.
 65. The method of claim54, wherein converting the attenuated fluence images to attenuated doseimages uses empirical calibration data.
 66. The method of claim 65,wherein empirical calibration data comprises 2D film dose images of thesame radiation beams used to generate the entrance beam images and acorrelation map between the entrance beam images and the 2D film doseimages.
 67. The method of claim 54, wherein converting the attenuatedfluence images to attenuated dose images comprises: acquiring dosemeasurements of radiation beams from the radiation source using adosimetric film, wherein the radiation beams are the same radiationbeams used to generate the entrance beam images; generating acorrelation map between the dose measurements and the entrance beamimages; and generating the attenuated dose images by applying thecorrelation map to the attenuated fluence images.
 68. The method ofclaim 67, wherein the acquired images of radiation beams are generatedfrom radiation measurements from a radiation detector comprising aplastic scintillating layer, and wherein acquiring dose measurements ofthe radiation beams comprises placing the dosimetric film in a mediumhaving similar properties to the plastic scintillating layer.
 69. Themethod of claim 68, wherein the plastic scintillating layer comprises alow-Z material having a thickness and the dosimetric film is placedinside a water or plastic phantom having the same thickness.
 70. Themethod of claim 55, wherein filtering the attenuated fluence imagescomprises using a scatter correction model to reduce radiation noise.71. The method of claim 54, wherein determining the dose distributioncomprises back projecting the exit beam image along each firing angle.72. The method of claim 54, wherein determining the dose distributionalong each firing angle of the radiation source uses patient CT imagesof attenuation coefficient numbers.
 73. The method of claim 72, whereincomputing a cumulative dose distribution map comprises convolution andsuperposition from a plurality of firing angles of a single platformposition.